The present invention relates to a two-dimensional array type of X-ray detector and a computerized tomography apparatus equipped with that X-ray detector.
The principal scan scheme in current X-ray computerized tomography apparatuses is the so-called third-generation type which collects projection data repeatedly while an X-ray tube and an X-ray detector revolve around a living body under examination. The principal X-ray detector is the so-called one-dimensional type in which a number of X-ray detecting elements are arranged in the channel direction.
In recent years, a scintillation type of X-ray detecting element has been put into practical use which consists of a combination of a scintillator and a photodiode. Having a good X-ray-to-electricity conversion efficiency and being compact and light, this type of detecting element surpasses an ionization chamber type of detector.
This compact, light X-ray detecting element has contributed greatly to the practical use of a two-dimensional array type of X-ray detector (also referred to as the multislice type of X-ray detector) in which a plurality of X-ray detecting elements are arrayed into a matrix. The two-dimensional array X-ray detector is manufactured roughly in two stages as shown in FIG. 1: in the first stage, a plurality of X-ray detecting elements corresponding in number to slices, for example, three X-ray detecting elements are arrayed along the Y direction (slice direction) to produce a module and, in the second stage, a plurality of modules corresponding in number to channels are arranged along the x direction (channel direction) so as to form a circular arc.
Here, a problem is that a module arrangement error arises in the second stage. That is, the X-ray detecting elements cannot be arranged accurately in a line in the channel direction. As a result, an artifact may be produced on a reconstructed image.
In addition, the scintillation type of X-ray detecting element has a property that the sensitivity drops abruptly in the vicinity of the edges of the scintillator as shown in FIG. 2 because the effective area of the photodiode is smaller than the plane of X-ray incidence of the scintillator. The half shadow of the collimator falls mainly on the edge portions of the scintillator. The half shadow greatly varies with variations in the geometrical position relationship of a collimator with the X-ray tube and/or the detector due to thermal expansion of parts of the X-ray tube. Thus, the sensitivity at the edges is very unstable, which may produce an artifact on the reconstructed image.
A method to solve the above problem is disclosed in Japanese Unexamined Patent Publication No. 5-184563. According to this method, a plurality of collimators are arranged at regular intervals along the slice direction between a body under examination and a two-dimensional X-ray detector. The collimator allows X-rays to be formed smaller than the dimension of the effective area of each X-ray detecting element in the slice direction. The resulting X-rays fall on only a portion 100 in the vicinity of the center of the effective area of each X-ray detecting element as shown in FIG. 1, thus allowing the effective area of the X-ray detecting element to have a margin. The module arrangement error is therefore allowed by the amount corresponding to that margin.
In Japanese Unexamined Patent Publication No. 9-224929, a collimator that is adjustable during scanning, called a dynamic pre-patient collimator, is provided between an X-ray tube and a body under examination, which allows the position of X-ray exposure to shift so as to reduce the effect of module arrangement errors.
However, the above-described two methods are accompanied by two problems that the half shadow is produced by the collimator and errors of collimator arrangement with respect to the X-ray tube and the X-ray detector are unavoidable. This means that the cause of the artifact is only changed from the module arrangement error to the two problems.